Ventilators are typically connected to patients via tubing and an interface. In invasive ventilation, the interface is usually a tube inserted in the patient's trachea (endotracheal tube; ET tube). A seal is usually maintained by inflating a cuff around the stem of the ET tube. In non-invasive ventilation (NIV), the interface is usually a mask or a helmet applied to the face/head around the patient's external airway (nose and/or mouth). A seal is maintained usually by applying appropriate tension to the straps holding the interface in place. In both invasive and non-invasive ventilation, the seal is at times imperfect resulting in air leaking from the ventilator circuit. In addition to these leaks at the interface, leaks can potentially arise from loose connections between different components of the tubing or defective tubing and, in the case of NIV, from deliberate leaks inserted in the tubing to act as a conduit for carbon dioxide removal. These different sources of leak display different relations between instantaneous pressure and leak flow. For example, with an orifice as the source of leak, leak flow increases as a function of the square root of pressure. With mask leaks, leak flow may increase linearly with flow or may increase disproportionately as pressure increases because the mask lifts away from the face, increasing the anatomical size of the leak. Thus, in practice, the relation between instantaneous pressure and leak flow is highly unpredictable.
Knowledge of the nature and magnitude of such leaks is important for the proper functioning of the ventilator. When the applied pressure is constant (for example, continuous positive airway pressure; CPAP) the leak is necessarily constant and can be readily estimated from the moving average of the flow signal. Since the amount of air entering the lungs during inspiration is roughly equal to the amount of air leaving the lungs during expiration, a moving average of flow rate over several breaths should be approximately zero in the absence of leaks. A positive average value indicates the presence of a leak and, because the leak is constant (since circuit pressure is constant), the magnitude of the moving average of flow rate reflects the magnitude of the constant leak.
The situation is quite different in ventilation modes in which the pressure applied to the patient is not constant (for example, bilevel support, proportional assist ventilation, volume cycled ventilation . . . etc). Here, the magnitude of the leak varies from time to time within each ventilator cycle depending on the instantaneous pressure in the circuit. Under these conditions, the moving average of flow rate continues to reflect the average leak but the value of the average leak no longer reflects the instantaneous leak at the different points in the cycle. Thus, average leak will overestimate the magnitude of leak during periods in which circuit pressure is below average, and vice versa. This could result in malfunction of the ventilator, particularly with respect to the times the ventilator is required to increase or decrease circuit pressure (triggering and cycling-off). For proper ventilator functioning it is necessary to know the relation between instantaneous circuit pressure and instantaneous leak. The present invention deals with an approach that determines said relation over the operating pressure range of the ventilation cycle.
Determination of the instantaneous relation between circuit pressure and leak has been problematic for the following reason. Flow rate is typically measured within the ventilator enclosure, usually at the points where the ventilating gas exits, and/or returns to the ventilator. The flow rate measured by the ventilator at any instant (Flow(i)TOTAL) is the sum of flow rate into or out of the patient (Flow(i)patient) and the flow rate going through the leak (Flow(i)leak). Thus:Flow(i)TOTAL=Flow(i)patient+Flow(i)leak 
As circuit pressure changes during the ventilator cycle Flow(i)TOTAL changes in part because Flow(i)patient changes and in part because Flow(i)leak also changes. In the course of a ventilator cycle Flow(i)patient changes in an unpredictable way, being subject not only to circuit pressure but also to the unknown time course of the patient's respiratory muscle pressure output (PMUS) and patient's respiratory mechanics. Because both components of Flow(i)TOTAL change unpredictably with instantaneous pressure, it is necessary to independently determine the relation between instantaneous pressure and leak flow rate in order to be able to estimate Flow(i)patient. It is knowledge of the patient's flow rate signal (Flow(i)patient) that is critical in the operation of the ventilator as it is used to trigger and cycle-off the ventilator as well as for monitoring adequacy of the assist and patient performance.
To the writer's knowledge, currently the most advanced method for estimating instantaneous leak during a variable pressure cycle is the one patented and used by Respironics (U.S. Pat. No. 5,803,065). Here, the total leak across a whole breath, or a plurality of breaths, is measured from the difference between the integral of Flow(i)TOTAL at the beginning of a breath and at the beginning of a subsequent breath, one or more breaths removed. This difference is assumed to reflect the volume of air that leaked out during the interval (ΔV). A certain mathematical function is then assumed for the relation between instantaneous pressure and instantaneous leak. The time course of pressure, processed according to this assumed function, is then integrated over the duration of the Interval. For example, if, as proposed in this prior art, the assumed function is [leak=K·P0.5], (i.e. leak is proportional to the square root of pressure), the integral of pressure, raised to the power 0.5 is calculated. The ratio ΔV/integral of P0.5 gives the value of K. Leak at any pressure can then be estimated from the leak equation according to the assumed function.
The above approach suffers from several drawbacks:
1) It is assumed that the volume difference between the beginning of a breath and the beginning of a subsequent breath (i.e. ΔV) is exclusively related to leaks. This is not true as the patient's inhaled and exhaled volumes can differ substantially on a breath-by-breath basis. Thus, ΔV across one breath incorporates an unknown component related to true differences in patient's lung volume between the beginning of one breath and the beginning of a subsequent breath (ΔVPATIENT). This problem is addressed in the Respironics technology by integrating the flow rate and the processed pressure over several breaths, on the valid assumption that differences between inhaled and exhaled volumes cancel out over a reasonable number of ventilator cycles (lung volume cannot indefinitely continue to rise or fall.
2) The above treatment assumes that the leak occurs over the entire pressure range, beginning with any pressure above zero (atmospheric). While this is true with built-in leaks, such as those inserted in the tubing during NIV, it is not true for unintentional leaks at the interface. Here, the seal may be air tight up to a certain pressure (i.e. no leak up to a threshold pressure (PTHRESH)). The leak would then be a function of pressure above this threshold [leak=K·ƒ(P−PTHRESH)]
3) Leaks often are derived from multiple sources. The treatment in this prior art assumes that one mathematical function can describe the overall leak even though such leak may be derived from multiple sources, each with its own function.
4) Most importantly, one needs to assume, a priori, the specific function that governs the relation between instantaneous pressure and instantaneous leak. In the prior art (U.S. Pat. No. 5,803,065) it was assumed that the function is a power function with the exponent being 0.5. Although this may accurately reflect the leak flow through a fixed orifice, naturally occurring leaks do not behave like fixed orifices. Rather, with the most common unintentional leaks, those at the interface (for example, insufficient tension in the mask straps), once pressure exceeds PTHRESH leak dimensions increase as pressure increases. Thus, not only is the driving pressure increasing, but the resistance to flow through the leak is also decreasing. In my observations, such leaks can follow a number of functions and, even assuming a power function, the exponent can be far from 0.5. Since it is not possible to know a priori what function these unintentional leaks will follow, it is not possible to implement the prior art for all possible types of leak and expect acceptable accuracy.
In summary, the correct relation between ΔV and a pressure-dependent leak is:ΔV−ΔVPATIENT=K·∫ƒ(P−PTHRESH)
Where, ΔVPATIENT is the change in patient's lung volume between the two measurement points, K is the constant of proportionality, PTHRESH is the pressure at which the leak begins, and ƒ is the specific function selected which may be any of several mathematical functions (for example, power, logarithmic, exponential, polynomial). In implementing the prior art it is not only necessary to know (or assume) the function type, but it is also necessary to specify the exponent of this function. Since neither the function, nor its exponent, nor PTHRESH can be known a priori, it can thus be seen that the approach described in the prior art [i.e. ΔV=K·∫P0.5] represents a great simplification and may be expected to provide erroneous instantaneous leak values in situations where the leak does not follow the behavior of fixed orifices.